The field of the invention is nuclear magnetic resonance imaging methods and systems. More particularly, the invention relates to the acquisition of an NMR image in which contrast is affected by magnetization transfer.
Any nucleus which possesses a magnetic moment attempts to align itself with the direction of the magnetic field in which it is located. In doing so, however, the nucleus precesses around this direction at a characteristic angular frequency (Larmor frequency) which is dependent on the strength of the magnetic field and on the properties of the specific nuclear species (the magnetogyric constant .gamma. of the nucleus). Nuclei which exhibit this phenomena are referred to herein as "spins".
When a substance such as human tissue is subjected to a uniform magnetic field (polarizing field B.sub.0), the individual magnetic moments of the spins in the tissue attempt to align with this polarizing field, but precess about it in random order at their characteristic Larmor frequency. A net magnetic moment M.sub.z is produced in the direction of the polarizing field, but the randomly oriented magnetic components in the orthogonal, or transverse, plane (x-y plane) cancel one another. If, however, the substance, or tissue, is subjected to a magnetic field (excitation field B.sub.1) which is in the x-y plane and which is near the Larmor frequency, the net aligned moment, M.sub.z, may be rotated, or "tipped" into the x-y plane to produce a net transverse magnetic moment M.sub.t, which is rotating, or spinning, in the x-y plane at the Larmor frequency.
The practical value of this phenomenon resides in the signal which is emitted by the excited spins after the excitation signal B.sub.1 is terminated. In simple systems the excited spins induce an oscillating sine wave signal in a receiving coil. The frequency of this signal is the Larmor frequency, and its initial amplitude, A.sub.0, is determined by the magnitude of the transverse magnetic moment M.sub.t. The amplitude, A, of the emission signal decays in an exponential fashion with time, t: EQU A=A.sub.0 e.sup.-t/T* 2
The decay constant 1/T*.sub.2 depends on the homogeneity of the magnetic field and on T.sub.2, which is referred to as the "spin-spin relaxation" constant, or the "transverse relaxation" constant. The T.sub.2 constant is inversely proportional to the exponential rate at which the aligned precession of the spins would dephase after removal of the excitation signal B.sub.1 in a perfectly homogeneous field. The practical value of the T.sub.2 constant is that spins associated with different tissue types have different T.sub.2 values and this can be exploited as a means of enhancing the contrast between such tissues. Of particular relevance to the present invention, however, is the fact that many spins are tightly coupled to each other and have T.sub.2 constants which are too short to acquire any useful signal from them with current NMR imaging pulse sequences. That is, these spins are excited and produce transverse magnetization, but they dephase quickly before the emission signal can be acquired.
Another important factor which contributes to the amplitude A of the NMR signal is referred to as the spin-lattice relaxation process which is characterized by the time constant T.sub.1. It describes the recovery of the net magnetic moment M to its equilibrium value along the axis of magnetic polarization (z). The T.sub.1 time constant is longer than T.sub.2, much longer in most substances of medical interest. As with the T.sub.2 constant, the difference in T.sub.1 between tissues can be exploited to provide image contrast.
When utilizing NMR to produce images, a technique is employed to obtain NMR signals from specific locations in the subject. Typically, the region which is to be imaged (region of interest) is scanned by a sequence of NMR measurement cycles which vary according to the particular localization method being used. The resulting set of received NMR signals are digitized and processed to reconstruct the image using one of many well known reconstruction techniques. To perform such a scan, it is, of course, necessary to elicit NMR signals from specific locations in the subject. This is accomplished by employing magnetic fields (G.sub.x, G.sub.y, and G.sub.z) which have the same direction as the polarizing field B.sub.0, but which have a gradient along the respective x, y and z axes. By controlling the strength of these gradients during each NMR cycle, the spatial distribution of spin excitation can be controlled and the location of the resulting NMR signals can be identified.
Conventional MRI images rely primarily on the signals produced by hydrogen nuclei associated with water and fat molecules in the human body. While the resonance frequencies of these two spin species is different (i.e. 210 Hz at 1.5 T), they both produce significant signal when the RF excitation field is tuned to either one (for example, 63.86 MHz for water at 1.5 T). The contrast in the resulting images is primarily determined by the density of each of these spin species throughout the slice or volume of interest, although this contrast can also be influenced, or "weighted" by taking advantage of the different T.sub.1 and T.sub.2 constants of these two spin species in various tissues of the body. Thus, for example, a pulse sequence might be designed to emphasize the signal from water spins having a relatively long T.sub.1 constant to enhance the contrast of tumor tissues in the image.
Image contrast can also be affected by selectively "saturating" certain spin species so that they are unable to emit any significant signal in the subsequently executed NMR pulse sequence. As described, for example, in co-pending U.S. patent application Ser. No. 07/902,634, such contrast preparation includes the application of an RF excitation pulse which may be specifically tuned to the resonance frequency of the spin species to be suppressed. The resulting transverse magnetization is dephased with spoiler gradients, and before the net magnet moment M.sub.z of the selected species can recover, the NMR pulse sequence is performed to acquire the desired signal. The spins which are saturated with the preparatory pulse contribute little to the acquired signal, thus affecting the image contrast of tissues containing such spins.
More recently, another mechanism has been discovered for modifying the contrast in MRI images. As described in U.S. Pat. No. 5,050,609, the contrast of MR images can be affected by selectively saturating spins which have a very short T.sub.2 constant and do not themselves contribute to the acquired signal Instead, this pool of "bound", or "restricted", spins may transfer magnetization to associated "free" spins and affect the NMR signals which are acquired from these free spins as well as the T.sub.1 and T.sub.2 values of these free spins. The amount of such magnetization transfer varies between tissue types, with the result that it provides yet another mechanism for weighting image contrast.
There are a number of methods used to produce magnetization transfer weighted MRI images. The initial method applied a continuous RF excitation signal to the region of interest which was tuned to a frequency different than that of water or fat. This is illustrated in FIG. 3 where curve 10 indicates the net magnetization of various short-T.sub.2 spin species having Larmor frequencies to either side of the "on resonance" frequency of the long-T.sub.2 water/fat spins. Dotted line 11 illustrates the very narrow band of short-T.sub.2 spins that are saturated with this method. While such "off resonance" continuous wave (CW) methods work, they present practical problems, such as the deposition of excessive RF power in human subjects and the need for separate RF amplifiers on commercially available MRI scanners. As a result, off resonance CW methods have not yet found widespread clinical use.
A number of pulsed magnetization transfer methods have also been proposed. Bob S. Hu et al. describe the application of a series of on resonance pulses which will saturate off resonance short T.sub.2 spins "Pulsed Saturation Transfer Contrast" Magn. Reson. Med. August 1992, 26(2), 231-240. Such binomial trains of pulses at the Larmor resonance of the desired long-T.sub.2 spin species (i.e. water/fat) will saturate spins outside a band of frequencies centered on resonance. Such a saturation curve for a 90.sub.x -90.sub.x -90.sub.x +90.sub.x (1111) binomial pulse sequence is shown by dashed line 12 in FIG. 3. It can be seen that a significant spectrum of short-T.sub.2 spins can be saturated with this method, thus making it more efficient than the CW method and enabling the RF power deposition in the patient to be lowered to acceptable levels. Practical considerations, however, also limit the applicability of this method. More specifically, curve 12 is the saturation that occurs in an ideal NMR scanner system with perfectly homogeneous magnetic fields. In practice, the inhomogeneities of the magnetic fields distort this saturation curve 12 such that long-T.sub.2 spins at and near resonance are also affected directly. In other words, the clean band of unsaturated spins around resonance is not achievable in practical NMR systems and the signals produced by spins associated with water and fat are reduced directly by such on resonance magnetization transfer pulses.
Yet another approach is to employ off resonance pulses which are shaped to avoid saturating spins near resonance. One such method described by Robert R. Edelman et al., "Improved Time-Of-Flight Angiography Of The Brain With Magnetization Transfer Contrast" Radiology August 1992, 184(2), 395-399, employs a pulse of RF at a frequency that is off resonance and has a Gaussian envelope. This produces the saturation curve 13 in FIG. 3. This method is robust in that it works well despite field inhomogeneities, but to achieve the ideal levels of saturation, the Gaussian pulse must be applied for 14 to 20 milliseconds. This long interval significantly increases the length of each pulse sequence and the length of the entire scan. A similar approach is proposed by Eric Outwater et al., "Magnetization Transfer Of Hepatic Lesion: Evaluation of a Novel Contrast Technique In The Abdomen" Radiology February 1992, 192(2) 535-540, where the off resonance RF pulse has a sinc (i.e. sin(x)/x) envelope. As with the Gaussian off resonance RF pulse, the sinc RF pulse must ideally be applied for 14 to 20 milliseconds to achieve the degree of short-T.sub.2 saturation needed for magnetization transfer contrast.